Pulmonary delivery of therapeutic agents can offer several advantages over other modes of delivery. These advantages include rapid onset, the convenience of patient self-administration, the potential for reduced drug side-effects, ease of delivery by inhalation, the elimination of needles, and the like. Inhalation therapy is capable of providing a drug delivery system that is easy to use in an inpatient or outpatient setting, results in very rapid onset of drug action, and produces minimal side effects.
Metered dose inhalers (MDIs) are used to deliver therapeutic agents to the respiratory tract. MDIs are generally suitable for administering therapeutic agents that can be formulated as solid respirable dry particles in a volatile liquid under pressure. Opening of a valve releases the suspension at relatively high velocity. The liquid then volatilizes, leaving behind a fast-moving aerosol of dry particles that contain the therapeutic agent.
Liquid aerosol delivery is one of the oldest forms of pulmonary drug delivery. Typically, liquid aerosols are created by an air jet nebulizer, which releases compressed air from a small orifice at high velocity, resulting in low pressure at the exit region due to the Bernoulli effect. See, e.g., U.S. Pat. No. 5,511,726. The low pressure is used to draw the fluid to be aerosolized out of a second tube. This fluid breaks into small droplets as it accelerates in the air stream. Disadvantages of this standard nebulizer design include relatively large primary liquid aerosol droplet size often requiring impaction of the primary droplet onto a baffle to generate secondary splash droplets of respirable sizes, lack of liquid aerosol droplet size uniformity, significant recirculation of the bulk drug solution, and low densities of small respirable liquid aerosol droplets in the inhaled air. In addition, a particular compound of interest may not be compatible with solvents typically used in nebulizer delivery systems.
Ultrasonic nebulizers use flat or concave piezoelectric disks submerged below a liquid reservoir to resonate the surface of the liquid reservoir, forming a liquid cone which sheds aerosol particles from its surface (U.S. 2006/0249144 and U.S. Pat. No. 5,551,416). Since no airflow is required in the aerosolization process, high aerosol concentrations can be achieved, however the piezoelectric components are relatively expensive to produce and are inefficient at aerosolizing suspensions, requiring active drug to be dissolved at low concentrations in water or saline solutions. Newer liquid aerosol technologies involve generating smaller and more uniform liquid respirable dry particles by passing the liquid to be aerosolized through micron-sized holes. See, e.g., U.S. Pat. No. 6,131,570; U.S. Pat. No. 5,724,957; and U.S. Pat. No. 6,098,620. Disadvantages of this technique include relatively expensive piezoelectric and fine mesh components as well as fouling of the holes from residual salts and from solid suspensions.
Dry powder inhalation has historically relied on lactose blending to allow for the dosing of particles that are small enough to be inhaled, but aren't dispersible enough on their own. This process is known to be inefficient and to not work for some drugs. Several groups have tried to improve on these shortcomings by developing dry powder inhaler (DPI) formulations that are respirable and dispersible and thus do not require lactose blending. Dry powder formulations for inhalation therapy are described in U.S. Pat. No. 5,993,805 to Sutton et al.; U.S. Pat. No. 69,216,527 to Platz et al.; WO 0000176 to Robinson et al.; WO 9916419 to Tarara et al.; WO 0000215 to Bot et al; U.S. Pat. No. 5,855,913 to Hanes et al.; and U.S. Pat. Nos. 6,136,295 and 5,874,064 to Edwards et al.
Broad clinical application of dry powder inhalation delivery has been limited by difficulties in generating dry powders of appropriate particle size, particle density, and dispersibility, in keeping the dry powder stored in a dry state, and in developing a convenient, hand-held device that effectively disperses the respirable dry particles to be inhaled in air. In addition, the particle size of dry powders for inhalation delivery is inherently limited by the fact that smaller respirable dry particles are harder to disperse in air. Dry powder formulations, while offering advantages over cumbersome liquid dosage forms and propellant-driven formulations, are prone to aggregation and low flowability, which considerably diminish dispersibility and the efficiency of dry powder-based inhalation therapies. For example, interparticular Van der Waals interactions and capillary condensation effects are known to contribute to aggregation of dry particles. Hickey, A. et al., “Factors Influencing the Dispersion of Dry Powders as Aerosols”, Pharmaceutical Technology, August, 1994.
To overcome interparticle adhesive forces, Batycky et al. in U.S. Pat. No. 7,182,961 teach production of so called “aerodynamically light respirable particles,” which have a volume median geometric diameter (VMGD) of greater than 5 microns (μm) as measured using a laser diffraction instrument such as HELOS (manufactured by Sympatec, Princeton, N.J.). See Batycky et al., column 7, lines 42-65. Another approach to improve dispersibility of respirable particles of average particle size of less than 10 μm, involves the addition of a water soluble polypeptide or addition of suitable excipients (including amino acid excipients such as leucine) in an amount of 50% to 99.9% by weight of the total composition. Eljamal et al., U.S. Pat. No. 6,582,729, column 4, lines 12-19 and column 5, line 55 to column 6, line 31. However, this approach reduces the amount of active agent that can be delivered using a fixed amount of powder. Therefore, an increased amount of dry powder is required to achieve the intended therapeutic results, for example, multiple inhalations and/or frequent administration may be required. Still other approaches involve the use of devices that apply mechanical forces, such as pressure from compressed gasses, to the small particles to disrupt interparticular adhesion during or just prior to administration. See, e.g., U.S. Pat. No. 7,601,336 to Lewis et al., U.S. Pat. No. 6,737,044 to Dickinson et al., U.S. Pat. No. 6,546,928 to Ashurst et al., or U.S. Pat. Applications 20090208582 to Johnston et al.
A further limitation that is shared by each of the above methods is that the aerosols produced typically include substantial quantities of inert carriers, solvents, emulsifiers, propellants, and other non-drug material. In general, the large quantities of non-drug material are required for effective formation of respirable dry particles small enough for alveolar delivery (e.g. less than 5 μm and preferably less than 3 μm). However, these amounts of non-drug material also serve to reduce the purity and amount of active drug substance that can be delivered. Thus, these methods remain substantially incapable of introducing large active drug dosages accurately to a patient for systemic delivery.
A thromboembolic event, such as myocardial infarction, deep venous thrombosis, pulmonary embolism, thrombotic stroke, etc., can present with certain symptoms that allow a patient or clinician to provide an initial therapy or treatment for the event. In some situations, an 81 mg, low dose, or baby aspirin or a regular aspirin (330 mg) may be orally administered in order to provide an initial treatment for the patient. In some cases, it has been recommended that upon first experiencing symptoms suspected of being due to a coronary event, patients are to chew and swallow two low dose aspirin tablets, a total dose of about 162 mg of aspirin.
There remains a need for providing novel formulations of non-steroidal anti-inflammatory drugs (“NSAIDs”), such as aspirin, that are suitable for pulmonary delivery.